1. Field of Invention
This invention relates to the field of tomography. More specifically, the present invention relates to a method for fiber-optically encoding attenuation data associated with detecting coincidences using a collimated source and a dedicated detector for improved measurement sensitivity.
2. Description of the Related Art
Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. Applications employing the PET technology for its sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
Using compounds such as 11C-labeled glucose, 18F-labeled glucose, 13N-labeled ammonia and 15O-labeled water, PET can be used to study such physiological phenomena as blood flow, tissue viability, and in vivo brain neuron activity. Positrons emitted by these neutron deficient compounds interact with free electrons in the body area of interest, resulting in the annihilation of the positron. This annihilation yields the simultaneous emission of a pair of photons (gamma rays) approximately 180xc2x0 (angular) apart. A compound having the desired physiological effect is administered to the patient, and the radiation resulting from annihilation is detected by a PET tomograph. After acquiring these annihilation xe2x80x9cevent pairsxe2x80x9d for a period of time, the isotope distribution in a cross section of the body can be reconstructed.
PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180xc2x0 angle of departure from the annihilation site, the location of the two detectors registering the xe2x80x9ceventxe2x80x9d define a chord passing through the location of the annihilation. By histogramming these lines of response (the chords), a xe2x80x9csinogramxe2x80x9d is produced that may be used by a process of back-projection to produce a three dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
In the traditional (2-D) acquisition of a modern PET tomograph, a collimator (usually tungsten) known as a septa is placed between the object within the field-of-view and the discrete axial rings of detectors. This septa limits the axial angle at which a gamma ray can impinge on a detector, typically limiting the number of axial rings of detectors that a given detector in a specific ring can form a coincidence with to a few rings toward the front of the tomograph from the given detector""s ring, the same ring that the detector is within, and a few rings toward the rear of the tomograph from the given detector""s ring.
Attenuation was first measured in PET by using a ring of positron emitting isotope surrounding the object to be measured. In this technique, the ratio between a transmission scan and a blank scan form the attenuation. The blank is measured by simply measuring the rate that gamma rays from positrons are detected by the detection system when no attenuating media is present. In the original scanners as described above as having septa, the septa are provided for collimating the gamma rays in an axial direction, but the rings allow for no transaxial collimation. The lack of collimation allow the acceptance of scattered events into the transmission measurement, resulting in an underestimate of the attenuation. To improve the transmission measurement, systems use rotating rod sources. These sources are disposed in parallel fashion to the axis of the scanner and are collimated in the axial direction by the septa. In the transaxial direction, the collimation may be provided electronically since the position of the source is known. However, the activity in the rod must be the same as that activity in the earlier ring source to provide the same count rate. With modern block detectors, the dead-time of the near block limits the activity in the rod.
A more recent advancement in PET acquisition is 3-D, in which the septa are removed, which allows a given detector to be in coincidence with detectors from all other detector rings. With the advent of three-dimensional reconstruction techniques, greater sensitivity to emission counts is possible if the septa are removed. As the septa represent a significant cost, there is also an economic incentive to exclude them from the system. However, with the absence of septa, the problems of both detector dead-time and scatter are magnified.
Since the position of a source with respect to the detector system can be known, there is no need to detect coincidences, thereby allowing the use of a source that emits single gamma rays. Only one detectorxe2x80x94the detector on the far side of the systemxe2x80x94is needed to make the transmission or blank measurements. Without the counting losses due to the dead-time of the near detector, the activity of the source may be increased resulting in an increase in count-rate and thus a better quality measurement. However, without axial collimation, the scatter included in the transmission scan causes an underestimate of the attenuation measurement. To decrease the possibility of scatter, the gamma rays from the source can be collimated with lead or tungsten to form a beam that illuminates only a narrow plane of detectors. Other gamma rays that would only contribute to background are eliminated. Since the directionality of single gamma rays cannot be determined, only a single point of activity illuminating a detector bank can be used. This requires increased levels of activity to meet the count-rate needed for an adequate quality measurement. Also, the scanning protocol is more efficient if the transmission measurement is performed after the patient has been injected with radioactivity. Even though a different isotope such as 137Cs which emits gamma rays with an energy of 662 keV can be used for the transmission scan, there is significant difficulty in distinguishing the transmission events from the emission events.
Another tomographic diagnostic system that is similar to PET is known as single photon emission computed tomography (SPECT). The distinction is that in SPECT, only a single photon from a nuclear decay within the patient is detected. Also, the line of response traveled by the photon is determined exclusively by detector collimation in SPECT, as opposed to the coincident detection of two collinear photons as in PET.
In computed axial tomography (CAT, or now also referred to as CT), an external x-ray source is caused to be passed around a patient. Detectors around the patient then respond to x-ray transmission through the patient to produce an image of an area of study. Unlike PET and SPECT, which are emission tomography techniques because they rely on detecting radiation emitted from the patient, CT is a transmission tomography technique which utilizes only a radiation source external to the patient.
The details of carrying out a PET study are given in numerous publications. Typically, the following references provide a background for PET. These are incorporated herein by reference for any of their teachings.
1. M. E. Phelps et al.: xe2x80x9cPositron Emission Tomography and Audiographyxe2x80x9d, Raven Press, 1986;
2. R. D. Evans: xe2x80x9cThe Atomic Nucleusxe2x80x9d, Kreiger, 1955;
3. J. C. Moyers: xe2x80x9cA High Performance Detector Electronics System for Positron Emission Tomographyxe2x80x9d, Masters Thesis, University of Tenn., Knoxville, Tenn., 1990;
4. U.S. Pat. No. 4,743,764 issued to M. E. Casey, et al, on May 10, 1988;
5. R. A. DeKemp et al.: xe2x80x9cAttenuation Correction in PET Using Single Photon Transmission Measurementxe2x80x9d, Med. Phys., vol. 21, 771-8, 1994;
6. S. R. Cherry et al.: xe2x80x9c3-D PET Using a Conventional Multislice Tomograph Without Septaxe2x80x9d, Jl. C.A.T., 15(4) 655-668.
7. J. S. Karp et al.: xe2x80x9cSingles Transmission in Volume-Imaging PET With a 137Cs Sourcexe2x80x9d, Phys. Med. Biol. Vol. 40, 929-944 (1995).
8. S. K. Yu et al.: xe2x80x9cSingle-Photon Transmission Measurements in Positron Tomography Using 137Csxe2x80x9d, Phys. Med. Biol. Vol. 40, 1255-1266 (1995).
9. G. F. Knoll: Radiation Detection and Measurement, John Wiley and Sons (1989).
10. S. R. Cherry et al.: xe2x80x9cOptical Fiber Readout of Scintillator Arrays using a Multi-Channel PMT: A High Resolution PET Detector for Animal Imagingxe2x80x9d, IEEE Transactions on Nuclear Science, Vol. 43, No. 3, 1932-1937 (June, 1996).
11. J. A. McIntyre et al.: xe2x80x9cConstruction of a Positron Emission Tomograph with 2.4 mm Detectorsxe2x80x9d, IEEE Transactions on Nuclear Science, Vol. 33, No. 1, 425-427 (February, 1986).
Both SPECT and CAT (or CT) systems are also well known to persons skilled in the art.
In order to achieve maximal quantitative measurement accuracy in tomography applications, an attenuation correction must be applied to the collected emission data. In a PET system, for example, this attenuation is dependent on both the total distance the two gamma rays must travel before striking the detector, and the density of the attenuating media in the path of travel. Depending on the location of the line of response within the patient""s body, large variations in attenuating media cross section and density have to be traversed. If not corrected, this attenuation causes unwanted spatial variations in the images that degrade the desired accuracy. As an example, for a cardiac study the attenuation is highest in the line of responses (LORs) passing through the width of the torso and arms, and attenuation is lowest in the LORs passing through from the front to the back of the chest.
Typically, the attenuation correction data in PET systems is produced by either: shape fitting and linear calculations using known attenuation constants, these being applicable to symmetric well-defined shapes such as the head and torso below the thorax (calculated attenuation); or through the measurement of the annihilation photon path""s attenuation using a separate transmission scan (measured attenuation). The use of calculated attenuation correction, which introduces no statistical noise into the emission data, can be automated for simple geometries such as the head, and is the most prominent method used for brain studies. However, complexities in the attenuation media geometry within the chest have prevented the application of calculated attenuation from being practical for studies within this region of the body. Accordingly, transmission scanning has been utilized.
The total attenuation of a beam along an LOR through an object is equal to the attenuation that occurs for the two photons from an annihilation. Thus, the emission attenuation along the path can be measured by placing a source of gamma rays on the LOR outside of the body and measuring attenuation through the body along this line. It has been the practice to accomplish this attenuation measurement by placing a cylindrical positron emitter xe2x80x9csheetxe2x80x9d within the PET tomograph""s field of view (FOV) but outside of the region (the object) to be measured. The ratio of an already acquired blank scan, where no object is placed in the FOV, to the acquired transmission scan is calculated. These data represent the desired measured attenuation factors, which may vary spatially. These data are then applied to the emission data after a transmission scan of the object to correct for the spatial variations in attenuation.
There are two types of transmitter source units conventionally utilized in PET transmission scan data collection, both of which form a xe2x80x9csheetxe2x80x9d of activity to surround the patient. These approaches include rotating dual photon rod sources (68Ge/Ga) with windowing and spiraling single photon point sources (137Cs) with collimation.
Rods with windowing are the most widely implemented mechanism for transmission measurements in PET. Advantages include noise rejection, physical simplicity (detectors optimized for emission event detection are also conscripted for transmission detection), and operational simplicity. L. R. Carroll et al., xe2x80x9cThe orbiting rod source: improving performance . . . ,xe2x80x9d Em. Com. Tom.:Current Trends, Soc. Of Nucl. Med., 1983, disclose windowing or masking for noise rejection. Operational simplicity is illustrated by S. R. Meikle et al., xe2x80x9cSimultaneous emission and transmission measurements . . . ,xe2x80x9d J. Nucl. Med., vol. 26, 1680-1688, 1995, who disclose the use of rods with windowing for simultaneous emission/transmission (SET) acquisitions.
A principal disadvantage of implementing rods with windowing for transmission measurements in PET is the high counting losses of the detectors nearest the rods. Typical of these detectors are: block decoding, as developed by M. E. Casey et al., xe2x80x9cA multicrystal two dimensional BGO detector . . . ,xe2x80x9d IEEE Trans. Nu. Sc., vol. 33, 460-463, 1986, in which photon scintillation by a single crystal affects counting losses of 31 or more nearby crystals; and electronic gated integration which also is a significant contributor to pulse processing time, which increases the likelihood of pileup. These two common features of current PET detectors, essential for emission photon detection, generate significantly high counting losses for detectors of photons nearest the transmission sources, or the near detectors, resulting in low counting statistics for practical duration transmission scans. Maximum noise equivalent rates for rod windowing of approximately 100 k events/sec are reported by W. F. Jones et al., xe2x80x9cOptimizing rod window width . . . ,xe2x80x9d IEEE Trans. Med. Im., vol. 14, 266-270, 1995.
As reported by R. A. DeKemp et al., single photon designs have been explored for overcoming the statistical limitations of dual photon rods. With no near detector, single photon offers an order of magnitude increase in gross transmission count rates. As reported by W. Jones et al., xe2x80x9cThe architectural impact of single photon . . . ,xe2x80x9d IEEE Nucl. Sci. Symp. Conf. Rec., 1026-1030, 1995, gross counts of several million events/sec are achievable. Unfortunately, high scatter fraction limits noise equivalent counts as reported by D. L. Bailey et al., xe2x80x9cStrategies for accurate . . . ,xe2x80x9d IEEE Nucl. Sci. Symp. Conf. Rec., 1997. While physical collimation reduces scatter content as disclosed by D. L. Bailey et al., xe2x80x9cA spiral CT approach . . . ,xe2x80x9d J. Nucl. Med., vol. 38, p. 113P (Abstract), 1997, the potential benefits from single photon are now viewed less optimistically. In the ART example reported by C. C. Watson et al., xe2x80x9cDesign and performance of single photon . . . ,xe2x80x9d IEEE Nucl. Sci. Symp. Conf. Rec., 1997, collimated single photon blank scan count rates are limited to 600 k to 2 M events/sec. Also for single photon design, innovative techniques make post-injection acquisitions practical and effective, but true SET remains less practical.
The ring source method significantly reduces the sensitivity of the tomograph due to the close source-proximity dead time effects of the source activity on all of the detectors. Further, removal of this assembly is either performed manually by facility personnel or by a complex automated mechanical assembly. Large, cumbersome, out of the FOV shielding is required for storage of the automated source when not in use, adding to the depth of the tomograph tunnel and, thus increasing incidence of patient claustrophobia. The second type of emitter, using rotating source(s) suffers from the above-mentioned problems and also, due to the shielding requirements, reduces the patient tunnel diameter, further increasing patient claustrophobia symptoms.
Both of the above automated source transportation methods suffer from high mechanical component cost and from low sensitivity. Due to the dead-time-induced reduction in tomograph sensitivity, lengthy acquisitions are required in order to achieve usable low noise transmission scan data.
In order to reduce costs in scintillator detector applications, multiplexing techniques based on the use of fiber optics are advantageous. Those disclosures made by Cherry et al. (Cherry), and McIntyre et al. (McIntyre), teach the use of fiber optics connected between the imaging detectors and multichannel photomultipliers (PMT""s). Cherry discloses the use of a multi-channel PMT in association with an 8xc3x978 array of bismuth germanate (BGO) crystals. As discussed by Cherry, a charge division readout board is used to convert the 64 signals into four position sensitive signals which determine the crystal interaction. In the earlier McIntyre article, the authors disclose the use of fiber optics coupled between the detectors and a number of multi-channel PMT""s. Specifically, McIntyre teaches the use of 288 PMT""s in association with 8,192 detectors, for reducing the number of required PMT""s by a factor of about 28.4.
In the McIntyre embodiment, eight detector rings are each divided into four quadrants. Each ring is comprised of sixteen concentric rings. The respective quadrants for the eight detector rings are grouped together for a total of 256 detectors per quadrant group. Sixteen xe2x80x9ccoarsexe2x80x9d fiber sets connect sixteen PMT""s to the 256 detectors, with sixteen detectors in one ring quadrant connected to one PMT. Similarly, sixteen xe2x80x9cfinexe2x80x9d fiber sets connect sixteen PMT""s to the 256 detectors, with corresponding detectors in each ring quadrant of a quadrant group being connected to one PMT. One PMT is connected to each ring quadrant. Thus, a total of 32 PMT""s are required for determining the particular detector xe2x80x9c"THgr"xe2x80x9d address within a quadrant. Similarly, 32 PMTs are required to determine the xe2x80x9crxe2x80x9d address, corresponding to which of the concentric rings in a particular ring the detector is disposed. Finally, eight PMTs are required to determine in which ring quadrant the detector is disposed. Thus, a total of 72 PMTs are required for each quadrant for a total of 288 PMTs in association with 8,192 detectors.
The present invention serves to detect coincident activity from a collimated point source. The present invention includes a detector dedicated to collecting attenuation data. The collimated point source and dedicated detector are positioned with respect to the tomography device such that only a selected strip of the imaging detector is illuminated such that events unrelated to the attenuation are eliminated. Fiber-optic encoding of the gamma radiation detectors is provided to minimize the required number of PMT""s.
The source of the present invention includes a collimator in which is disposed a point source. An opening is defined by the collimator for exposing a selected portion of the imaging detectors of the tomograph device. Positioned behind the point source, relative to the imaging detectors, is an attenuation detector dedicated to collecting attenuation data. Because the attenuation detector is dedicated to the attenuation measurement, the requirements for the attenuation detector are different from those for the imaging detector. For instance, it is not required that the attenuation detector be able to accurately determine the energy or spatial position of events within the detector, as is necessary for standard imaging detectors. It is therefore possible to design such an attenuation detector with much less dead time, and much higher count rate performance, than a standard imaging detector. The improved count rate performance of the attenuation detector enables significant reduction of statistical noise in the attenuation correction measurement. The attenuation detector and collimator are designed to illuminate only a strip of the imaging detector, and the corresponding aperture of the attenuation detector, thereby eliminating events not of interest in the attenuation measurement. This also reduces dead time of the system and improves the count rate performance for events of interest.
A source of the present invention is disposed opposite each bank of imaging detectors of a dual head camera. Each source contains four point sources arranged along the axial extent. The sources and the associated collimators are positioned to the side of each head at a slight angle relative to the respective head. The sources and detectors are fixed relative to the imaging heads. In order to obtain full coverage of the field of view (FOV) in the same manner as for an emission scan, the heads and sources are rotated about the center of the camera. In one embodiment, a two-dimensional array of detectors is associated with each head. The detectors are grouped in threes, with the first of each group being aligned, the second of each group being aligned with each other and transaxially positioned relative to the first, and the third of each group being aligned with each other and transaxially positioned relative to the first and second. This disposition of the detectors substantially reduces crosstalk between successive pairs.
The present invention further provides an arrangement of fiber optics interconnected between a plurality of dedicated gamma radiation detectors and a lesser number of photomultiplier tubes. The gamma radiation detectors are each provided for dedicated detection of 511 keV gamma radiation from one of a plurality of point sources disposed in a collimator. The arrangement of fiber optics is designed such that the address of a particular gamma radiation detector is readily discernable while minimizing the number of PMT""s required to process data accumulated by the gamma radiation detectors.
In one embodiment, the attenuation point sources are disposed in a two-dimensional array having three rows, thus yielding three FOV fans defined between the transmission sources and the planar detector arrays. By moving neighboring source/crystal pairs transaxially, sufficient lead and distance is added to effectively minimize gamma crosstalk.